High energy ionizing radiation such as a gamma ray cannot be detected directly; it must be converted into an electrical pulse. One common method for creating an electrical pulse when ionizing radiation is present is to first absorb the ionizing radiation in a scintillator. In response, the scintillator then produces a flash of light which is converted into an electrical signal by a photodetector.
Although the physical configuration of the scintillator/photodetector combination varies from application to application, the underlying principle remains constant. The scintillator will vary in size, but must be thick enough to stop the incident radiation and large enough to cover the desired area. The exact type of photodetector that the scintillator is coupled to varies but the photodetector must produce an electrical signal large enough to be observable above background noise.
The effectiveness of this method of radiation detection is primarily limited by the scintillating material. Only a few materials are known to scintillate, and the scintillation properties of these materials vary. An ideal scintillator would have a high density, a short decay time constant (i.e. the photons are emitted as soon as possible after the radiation interacts in the scintillator), and a large light output that is essentially proportional to the amount of radiation deposited in the scintillator. High density is desirable in order to stop the ionizing radiation in as short a distance as possible, a short decay time is desirable in order to measure the time of interaction accurately, and a high light output is desirable in order to make it easier for the photodetector to convert the light into an electrical pulse. In addition, the scintillating material should not have unpleasant chemical or material properties, such as toxicity, hygroscopy, or extreme reactivity.
The photodetector that converts the scintillation light into an electrical pulse is usually either a photomultiplier tube or a solid-state (Si, GaAs, HgI, etc.) photodiode. A photomultiplier tube (PMT) is usually employed when very small amounts of radiation are to be detected, as a PMT typically converts each photon into approximately 1 million electrons. Photodiodes are usually used when large amounts of radiation are to be detected, as a photodiode typically converts each photon into a single electron. These photodetectors can either be coupled to the scintillator either singly or in position sensitive arrays.
FIG. 1 diagrammatically demonstrates a single embodiment of a high energy radiation detector contemplated in the present invention. The device 10 is in the form of a hand held radiation monitor and is similar to a standard Geiger counter, but will be more sensitive to gamma radiation. A scintillator crystal 11 is optically coupled to a photomultiplier tube 12 and this assembly is encased in an opaque material 13 in order to shield ambient light. A cable 14 connects the scintillator/photodetector assembly to a small box 15 containing a battery operated power supply for the photomultiplier tube, counting electronics, and a numerical display. When ionizing radiation 17 from a source 16 impinges on the scintillator crystal, it emits photons which are converted into an electrical pulse by the photomultiplier tube 12 and the rate at which these pulses arrive is determined by the electronic circuit in the box 15 and displayed accordingly. As this rate is proportional to the amount of radiation present, the device can be used to monitor radiation.
A more complicated embodiment of a radiation detector is a radioisotope camera. Radioisotope cameras have considerable important uses in medical diagnosis and research. They have a significant advantage over the use of X-ray techniques in that they are sensitive to minute amounts of certain radioactive tracer compounds. In use, a small quantity of a gamma-ray emitting radioactive substance is injected into a patient. The choice of radioactive isotope depends on its half-life, activity, dose rate, and many other factors. Certain isotopes are highly specific, and tend to concentrate in certain organs of the body. This selective accumulation permits visualization of the biological function of almost every organ of the human body. Although the resolution of radioisotope cameras is not equal to that obtained in X-ray radiographs, the size, shape, position and function of the organs can be determined, and often lesions can be located in them.
PET is a medical imaging technique in which a radioactively labeled substance is administered to a patient and then traced within the patient's body by means of an instrument that detects the decay of the isotope. In the positron emission tomographic process a chemical tracer compound having a desired biological activity or affinity for a particular organ is labeled with a radioactive isotope that decays by emitting a positron (positive electron). The emitted positron loses most of its kinetic energy after traveling only a few millimeters in living tissue. It is then highly susceptible to interaction with an electron, an event that annihilates both particles. The mass of the two particles is converted into 1.02 million electron volts (1.02 MeV) of energy, divided equally between two 511 keV photons (gamma rays). The two photons are emitted simultaneously and travel in opposite directions. The two photons penetrate the surrounding tissue, exit the patient's body, and are absorbed and recorded by radiation detectors geometrically arranged to record the simultaneous emission event.
The instrumentation used in PET is a camera generally in a geometric arrangement consisting of an array of hundreds of scintillation elements in a ring surrounding a radioactive emitting subject. Assuming no scattering, the simultaneously emitted gamma-rays strike two scintillation elements perpendicularly disposed 180 degrees from each other in the scintillation ring array. The scintillation flashes will be detected by photodetectors coupled to the scintillators. Row and column amplifiers and a position logic circuit connected to the scintillator array will then distinguish the electrical pulse generated by the activated scintillation elements. A state of the art PET camera is demonstrated in U.S. Pat. No. 4,672,207 to DERENZO. This patent discloses a radioisotope camera in which an array of scintillation crystals optically coupled to photodetectors arranged in rows and columns and adapted to be exposed to a radioactive emitting subject. The array is further coupled to means for reconstructing the image created by the radioactive emissions from the subject.
Biological activity within an organ under investigation can be assessed by tracing the source of the radiation emitted from the patient's body to the scintillators of the PET camera. The source of the radiation can be accurately estimated by linking each scintillator crystal element with several other scintillators on the opposite side of the array and registering a signal only if two detectors sense 511 keV photons coincidentally (typically 10 nsec). When a coincidence is registered, an annihilation is recorded along a line connecting the two scintillators. In this manner, a circumferential array of photodetectors can establish the sources of all coincident pairs of photons that originate within a volume defined by straight lines joining paired detectors. A computer program reconstructs the spatial distribution of the decaying isotopes within the patient. With suitable interpretation, PET images provide a noninvasive, regional assessment of many biochemical processes associated with human organs.
The value of PET as a clinical imaging technique is in large measure dependent upon the performance of the radiation detection elements. The typical PET camera comprises an array of radiation detectors consisting of scintillator crystals coupled to photomultiplier tubes (PMT's). When a photon strikes a detector, it produces light in one of the scintillator crystals that is then sensed by the PMT, which registers the event by passing an electronic signal to the reconstruction processing circuitry.
As pointed out earlier, the scintillator crystals used in a general purpose detector or a PET camera must have certain properties, among which are (1) good stopping power, (2) high light yield, and (3) fast decay time.
In a PET application, stopping power is the ability to stop the 511 keV photons in as little material as possible so as to reduce the overall size of the scintillator, which reduces the cost and improves spatial accuracy. Stopping power is typically expressed as the linear attenuation coefficient (tau) having units of inverse cm.sup.-1. After a photon beam has traveled a distance "x" in a crystal, the fraction of photons that have not been stopped by the crystal is calculated as follows: EQU fraction of unstopped photons=e.sup.(-tau * x).
Therefore, after traveling a distance of 1/tau (the "absorption length"), approximately 37% of the photons will not have been stopped: 63% will have been stopped. Likewise, 63% of the remaining photons will have been stopped after traveling an additional distance of 1/tau. For PET, one wants 1/tau to be as small as possible so that the photodetector is as compact as possible.
Light yield is also an important property of scintillators contemplated for use in PET. Light yield is sometimes referred to as light output or relative scintillation output, and is typically expressed as the percentage of light output from a crystal exposed to `standard` scintillation 511 keV photons relative to the light output from a crystal, thallium-doped sodium iodide, NaI(Tl) under the same conditions. Accordingly, the light yield for NaI(Tl) is defined as 100.
A third important property of scintillators in PET applications is decay time. Scintillation decay time, sometimes referred to as the time constant or decay constant, is a measure of the duration of the light pulse emitted by a scintillator, and is typically expressed in units of nanoseconds (nsec). As noted above, in PET, the source of biological activity within an organ under investigation is determined by tracing the source of coincident photons emitted from the patient's body to the photodetectors. When two 511 keV photons are detected at the same time by a pair of scintillators, the source of the photons is known to lie along the linear path connecting the two scintillators. In general, only a fraction of the detected photons are in coincidence and thus used in the reconstruction analysis. Moreover, many false coincidences are registered because the finite decay time associated with each scintillator may cause it to emit light at the same time as another scintillator when in fact the photons inducing the light did not come from the same positron annihilation. For example, a photon arrived at one photodetector may produce a flash of light that does not decay, i.e. "turn off", until after a later photon, from a different positron annihilation, produces a flash of light in a detector on the side opposite the first detector. In this instance, the flashes would overlap, and the photodetectors would register them as in coincidence. Thus, scintillator materials with long decay constants have an inherent problem in detecting coincident photons.
In addition to the problem of false coincidences, the positron emitting tracer compounds themselves typically have very short half-lives. In fact, most medical facilities performing PET also operate on-site accelerators to produce the short-lived radioactively labeled tracer compounds. Because of the short half-lives of these compounds, data on the occurrence of coincident photons needs to be gathered at as high a rate as possible. As noted above, the majority of the detected photons are not in coincidence, i.e., they are from sources outside the plane of the detector array. Consequently, if a scintillator's decay constant is short, then more of its time will be available for the detection of coincident photons.
In addition to the three important properties discussed above, scintillator crystals for PET should be easy to handle. For example, certain known scintillators are very hygroscopic, i.e., they react with moisture, making it necessary to very tightly encapsulate them to allow their use as scintillators in PET. These hygroscopic scintillators are expensive and difficult to use.
Known scintillator materials include (1) plastic scintillators, (2) thallium-doped sodium iodide (NaI(Tl)), (3) cesium fluoride (CsF), (4) bismuth germanate (Bi.sub.4 Ge.sub.3 O.sub.12, also referred to as "BGO"), (5) cerium fluoride (CeF.sub.3), and (6) barium fluoride (BaF.sub.2). Of these five scintillators, only two, BGO and BaF.sub.2, are used routinely for PET.
Plastic scintillators, typically composed of polystyrene doped with a wavelength-shifting additive, are commercially available under such tradenames as PILOT U and NE 111. Upon excitation with a 511 keV photon, plastic scintillators emit a light pulse having a very fast decay constant of approximately 1.5 nsec and light output proportional to the energy of the incident photon. The main disadvantage of plastic scintillators is their low density (approximately 1.1 to 1.2 g/cm.sup.3) due to the light atoms (hydrogen and carbon) that make up the molecules of the material. Because of their low density, plastic scintillators have poor stopping power, and are therefore poorly suited for use in PET.
NaI(Tl), thallium-doped sodium iodide, has the highest light output of the six scintillators listed above. NaI(Tl) also has an intermediate stopping power (1/tau=3.0 cm at 511 keV). However, NaI(Tl) has a long decay constant (250 nsec), a significant disadvantage for use in PET. NaI(Tl) has an additional disadvantage in that it is highly hygroscopic, making it extremely difficult to handle in that it must be tightly encapsulated in bulky cans.
CsF, cesium fluoride, has an advantage over plastic scintillators because of its relatively high density (4.61 g/cm.sup.3) and consequent stopping power. However, the light output and decay constant of CsF are inferior to those of plastic scintillators. CsF is also highly hygroscopic, well above NaI(Tl) which, as noted above, makes it expensive and difficult to handle.
BGO has the highest density (7.13 g/cm) of the five known scintillator materials noted above. Its stopping power is the best of the five materials (1/tau=1.1 cm at 511 keV). As a result, BGO is best able to absorb 511 keV photons efficiently in small crystals. However, BGO's very long decay constant (300 nsec), longer even than NaI(Tl), is at a significant disadvantage for use in PET.
Cerium fluoride, CeF.sub.3, a recently discovered scintillator, has a relatively high density (6.2 g/cm.sup.3) and correspondingly high stopping power. However, its decay time is moderately long (27 ns) and its light output is low.
The use of BaF.sub.2 as a scintillator material is described in Allemand et al. U.S. Pat. No. 4,510,394. BaF.sub.2 emits light having two components: 75% is emitted with a `slow` decay constant of approximately 620 nsec and 25% with a `fast` decay constant of approximately 0.6 nsec. BaF.sub.2 has a light yield of approximately 16% that of NaI(Tl) and about half the stopping power of BGO (1/tau=2.3 cm at 511 keV). Unlike CsF and NaI(Tl), BaF.sub.2 is not hygroscopic.
The fast component of BaF.sub.2 emits light in the ultraviolet region of the spectrum. Glass photomultiplier tubes are not transparent to ultraviolet light, so a quartz photomultiplier tube must be used instead to detect the fast component of BaF.sub.2. Since quartz photomultiplier tubes are substantially more expensive than glass, one would prefer to avoid using BaF.sub.2, if possible, in favor of using a scintillator that can be detected by a glass photomultiplier tube. The fast component gives BaF.sub.2 very good timing resolution, but the slow component limits its high rate capabilities. In other words, it takes longer when using BaF.sub.2 to get ready for the next event.
Of the best known scintillator materials, BGO has the best stopping power, NaI(Tl) has the best light yield, and BaF.sub.2 has the best timing resolution. However, as noted above, each of these materials have significant shortcomings which hinder their performance as scintillators for PET: BGO has a very long decay constant, NaI(Tl) has a very long decay constant comparatively low density and is hygroscopic, and BaF.sub.2 has a long decay constant and requires expensive photomultiplier tubes.